Microfabricated nanopore device for sustained release of therapeutic agent

ABSTRACT

A drug delivery device that includes a capsule for implantation into the body; the capsule further includes a reservoir for containing a substance such as a therapeutic agent, at least one port for allowing the substance to diffuse from or otherwise exit the reservoir, and a nanopore membrane in communication with the capsule at or near the exit port for controlling the rate of diffusion of the substance from the exit port. The device also includes an optional screen for providing structural stability to the nanopore membrane and for keeping the pores of the nanopore membrane clear. One embodiment of the drug delivery device includes an osmotic engine internal to the device for creating fluid flow through the device.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.13/098,067, filed Apr. 29, 2011, now allowed, which is a continuation ofU.S. application Ser. No. 11/530,729, filed Sep. 11, 2006, now U.S. Pat.No. 7,955,614, which is a continuation of U.S. application Ser. No.10/243,787, filed Sep. 13, 2002, now abandoned, which claims the benefitof U.S. Provisional Application No. 60/322,160 filed on Sep. 14, 2001,entitled “Microfabricated Nanopore Device for Sustained Release ofTherapeutic Agent” and U.S. Provisional Application No. 60/371,290 filedon Apr. 9, 2002 entitled “Microfabricated Nanopore Device for SustainedRelease of Therapeutic Agent,” the disclosures of which are fullyincorporated by reference in their entirety as if fully rewrittenherein.

TECHNICAL FIELD OF THE INVENTION

The present invention relates generally to drug delivery devicesdesigned to be implanted in the body and specifically to an implantabledrug delivery device that utilizes a nanopore membrane fabricated toinclude arrays of channels having precise geometry for controlling therelease of a therapeutic agent into the body.

BACKGROUND OF THE INVENTION

The biotechnology industry has created a number of successful biologicaldrugs for treating chronic medical conditions. Such drugs includeincluding alpha epoetin (Procrit®, Epogen®) for treatment of chronicanemia associated with cancer chemotherapy, granulocyte colonystimulating factor (Neupogen®) for treatment of neutropenia associatedwith cancer chemotherapy; alpha interferon (Intron® A, Roferon® andInfergen®) for treatment of chronic hepatitis; and, beta Interferon(Avonex®) for treatment of relapsing multiple sclerosis.

Although these products provide important benefits to patients andgenerate sales measured in billions of dollars, these and most otherbiological drugs have two major limitations: (i) due to their largemolecular weight (e.g., >10,000 Daltons) and fragility, such drugscannot be delivered to the patient by the oral route, and thus,injection is the only method of administration; and (ii) their typicallyshort half-life results in the drugs being quickly cleared from thebody, and therefore, they must be administered to the patient frequently(e.g., daily or three times per week).

In the hospital setting, intravenous administration is usually a safeand reliable method for administering biological drugs. In medicalsituations where hospitalization or physician visits are not necessary,patients often self-administer biological drugs by subcutaneous orintramuscular injection several times a week over the course of therapy.However, this type of therapy is generally associated with pain at thesite of injection, injection site reactions, infections, and lack ofcompliance with dosing schedule.

Sustained release implants or drug depots provide a potential solutionto the medical need for delivering biological drugs for chronicconditions. Sustained release implants have the potential to eliminatecompliance as a concern because they provide the physician with theassurance that the drug is being delivered and the patient with thefreedom to go about their normal daily activities. Currently, two basictechnologies have been developed to address the medical need forsustained release of chronically administered injectable drugs:injectable erodable polymer depots designed to act for several weeks,and implantable devices capable of delivering potent drugs for up to oneyear. While effective in certain cases, these prior art devices aresubject to important limitations.

Sustained release depot formulations that employ polymer depotstypically exhibit an initial “burst effect” resulting in the release ofup to 90% of the encapsulated drug in the first few days afterimplantation. Following injection of the device, plasma levels quicklypeak and then decline to near constant levels. This characteristic ofdepots makes them unsuitable for sustained release of a drug over time,where a more constant rate of delivery is desired.

Other prior art implantable devices utilize a semi-permeable membrane tocause osmotic tablets to slowly swell as they absorb water. The swellingtablets push a piston that forces drug from a reservoir out of a smallopening. Such devices are capable of sustained release over longerperiods of time; however, the number of drugs that are compatible withsuch devices is limited due to the construction of the device. Thus,only highly potent drugs such as certain hormones can be successfullyused with these prior art devices.

Thus, although current technologies provide important advantages overtraditional daily injections, a need currently exists for implantablesystems that are more flexible with respect to the types, size,stability and solubility properties of drugs themselves, and withrespect to the drug delivery patterns achievable using such a device,i.e., avoidance of “burst” effects and achieving a more constant rate ofdrug delivery.

SUMMARY OF THE INVENTION

These and other limitations of the prior art are overcome by the presentinvention which provides a drug delivery device that includes a capsulefor implantation into the body. This capsule effectively separates thedrug or other material that is encapsulated within the device from theenvironment external to the device. The capsule further includes areservoir for containing a substance such as a therapeutic agent, atleast one port for allowing the substance to diffuse from or otherwiseexit the reservoir; and a nanopore membrane in communication with thecapsule at or near the exit port for controlling the rate of diffusionof the substance from the exit port. The device also includes a screenfor providing structural support to the nanopore membrane and forkeeping the pores of the nanopore membrane clear.

The capsule includes a nanopore membrane consisting of an array ofparallel channels with precise dimensions typically in the 4 to 100 nmrange separating the internal reservoir from the external medium. Byprecisely tailoring the pores of the membrane to the moleculardimensions of the drug, such nanopore membranes serve to control thediffusion kinetics of the therapeutic agent from the reservoir at amolecular level. Presumably, the rate of diffusion is related to thegeometry of the channels which physically constrain the random molecularmotion of drug solute molecules in at least one dimension. Thus, therate of diffusion is slowed and controlled at a nearly constant level asa function of pore size and is less dependent on the concentrationgradient. The device is capable of zero-order release of a therapeuticagent after implantation over prolonged periods (weeks to months, oreven years). Controlling the dosage delivered to the patient over timeis also possible with the present invention.

One embodiment of the drug delivery device includes an osmotic enginefor creating a flow of fluid into and through the device. The osmoticengine further includes a semi-permeable membrane incorporated into aportion of the wall of the capsule, and an osmotically active agentadmixed with the therapeutic agent. The osmotically active agent is ofsufficient molecular weight as to be restricted from passing throughboth the semi-permeable membrane and the nanopore membrane. The osmoticengine also includes a net flux of water entering the reservoir from theexternal medium through the semi-permeable membrane and exiting thedevice through the nanopore membrane.

Drugs therapies that meet the following criteria are good candidates forformulation in the present implant device: (i) those that normallyinjected intravenously or subcutaneously; (ii) those that requirefrequent administration over a prolonged period of time (e.g., 2-3 timesper week for more than two weeks); (iii) those involving treatment of aserious condition where non-compliance to the prescribed treatmentregimen will have serious consequences; (iv) those utilizingsufficiently potent drugs so that the cumulative dose for the period oftreatment can fit within a small reservoir; and/or (v) those using adrug with adequate stability to withstand exposure to body temperaturefor duration of therapy.

Further advantages of the present invention will become apparent tothose of ordinary skill in the art upon reading and understanding thefollowing detailed description of the preferred embodiments.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings and figures, which are incorporated in andform a part of the specification, schematically illustrate preferredembodiments of the invention and, together with the general descriptiongiven above and detailed description of the preferred embodiments andexamples given below, serve to explain the principles of the invention.

FIG. 1 is a graphic representation of the basic steps in themicrofabrication of the nanopore membrane of the present invention.

FIG. 2 is a graphic illustration of the basic structural features of thenanopore membrane; both a top view and a side view are presented.

FIG. 3 a is a graphic illustration of one embodiment of the drugdelivery device of the present invention showing a capsule comprising asegment of thin-walled titanium alloy tubing, sealed at each end withpolymer end-caps, and a nanopore membrane fixed to a milled out portionof the capsule.

FIG. 3 b is a graphic illustration of another embodiment of the drugdelivery device of the present invention providing an exploded view ofthe device and its component parts.

FIG. 3 c is a graphic illustration of another embodiment of the drugdelivery device of the present invention including adiffusion-facilitating osmotic engine. This embodiment includes asemipermeable membrane and osmagent, which combine to provide passage ofwater from the external medium, through the reservoir of the device andout the nanopore membrane channel array (indicated by the dotted line).

FIG. 4 is a graph showing the relative rates of glucose diffusionthrough nanopore membranes with pore widths ranging from 7 nm to 27 nm.

FIG. 5 plots the rates of glucose diffusion through nanopore membranesof sizes 7 nm, 13 nm and 49 nm, as a function of initial glucoseconcentration.

FIG. 6 is a graph showing the cumulative release of ¹²⁵-I labeledalbumin from two devices similar to the one illustrated in FIG. 3 a. Thedevices are outfitted with nanopore membranes of two pore widths, 26 nmand 13 nm. The dashed lines represent the diffusion rates that arepredicted from Fick's laws of diffusion.

FIG. 7 is a plot showing blood levels of ¹²⁵I-albumin in groups of ratsgiven either a subcutaneous injection of 125 micrograms of ¹²⁵-I labeledalbumin (squares), or implanted with nanopore membrane drug deliverydevices designed to release the albumin at rates of 7 μg/day (circles)and 15 μg/day (triangles).

FIG. 8 shows the in vitro diffusion of ¹²⁵-I labeled lysozyme as afunction of time through nanopore membranes with pore widths of 7 nm(diamonds), 13 nm (squares) and 49 nm (triangles).

FIG. 9 is graph showing the blood levels over a 50-day period of ¹²⁵-Ilabeled lysozyme in groups of rats after subcutaneous injection of 80 μgof lysozyme (dashed line) or after subcutaneous implantation of ananopore membrane device designed to release 17 μg/day of lysozyme.

DETAILED DESCRIPTION OF THE INVENTION I. Drug Delivery Implant Device

The present invention is designed as a small medical implant forsustained, constant release of potent drugs used to treat chronicconditions. An exemplary embodiment is substantially cylindrical inshape and has a diameter of about 4 to 8 millimeters (about 0.125 to0.25 inches) and a length of about 40 to 80 millimeters (about 1.5 to 3inches). Preferably, the device is implanted in a manner similar toother medical implants; following administration of a local anesthetic,a physician or other medical professional makes a small (about 5millimeter) incision in the skin of the upper arm, forearm or abdomen ofthe patient. Using a sterile implanting tool, the physician inserts thedevice under the skin of the patient (i.e., subcutaneously) and thencovers the incision with a bandage. At the end of the treatment period,the physician removes the implant using a similar procedure.

The drug release device of the present invention is designed to slowlyrelease the encapsulated drug at a nearly constant rate to mimic a slowinfusion, so that the patient will have therapeutic levels of the drugin his/her body for the entire course of therapy. The drug reservoir ofdevice contains a highly concentrated form of the drug either as asaturated solution, dry powder or concentrated suspension to minimizethe size of the device required to hold the cumulative dose required foran extended period of treatment (e.g., several weeks to 6 months).

As shown in FIGS. 3 a-c, an exemplary embodiment of the presentinvention provides a drug delivery device 10 which includes anencasement or capsule 12 for implantation into the body. Capsule 12further includes reservoir 22 for containing a substance, such as atherapeutic agent, and at least one port 24 for allowing the substanceto diffuse from or otherwise exit the reservoir, and a nanopore membrane26 attached to, mounted on, or otherwise in communication with capsule12 and exit port 24 for controlling the rate of diffusion of thesubstance from exit port 24. In this embodiment, membrane 26 serves asthe only diffusion-limiting communication between the reservoir and theexternal medium.

Optionally, the drug delivery device may include a screen 28 forproviding structural stability to the nanopore membrane withoutaffecting the rate of release of a therapeutic agent. This screen istypically made from a porous polymer or other material and maycompletely or partially surround the capsule, or may be positioned ontop of the membrane, on the bottom of the membrane, or both on the topand bottom of the membrane. Screen 28 also prevents fouling/blockage orcellular infiltration of the pores of nanopore membrane 26. In oneembodiment the entire capsule is coating with a porous polymericmaterial to protect the device and provide a biocompatible interface.

An exemplary embodiment of the capsule of the drug delivery device ofthe present invention is an impermeable, non-deformable, biocompatiblecapsule which further includes a first open end 14; a second open end18; a first polymeric end cap 16 for closing first open end 14; and apolymeric second end cap 20 for closing second open end 18. The openends are useful for filling reservoir 22 with a therapeutic agent orother substance. In one embodiment, the end caps are tapered such thatwhen attached to each open end of the capsule, both ends of the capsuleare bullet-shaped. Capsule 12 may be substantially cylindrical incross-section or substantially elliptical in cross-section. Asemi-triangular anti-rotation device may be attached to each end of thecapsule, or anywhere along the length of the capsule, for preventingside-to-side rotation of the capsule following implantation of the drugdelivery device. Capsule 12 may be manufactured from titanium alloy,surgical grade stainless steel, or a polymeric material.

The drug delivery device of the present invention is designed forimplantation in the body subcutaneously, in a defined anatomicalcompartment of the body, at a pathological site, or at combination ofsites. Suitable sites include the peritoneal cavity, brain, pluralcavity, surgical site, pathological site, pericardium, inter articularspace, eye, and/or subarachnoid space (C SF).

In an exemplary embodiment, capsule 12 is typically made of segments ofstandard thin-walled stainless steel or titanium alloy tubing in which aflat area has been milled to serve as the seat for nanopore membrane 26.Nanopore membrane 26 is mounted over this area of the device underprotective screen 28. In one embodiment, drug delivery device 10 isabout 4 to 10 mm in diameter and about 45-100 mm in length with aninternal volume of approximately about 250 μL to several millilitersdepending on the thickness of the tubing.

In another embodiment, a small area of the capsule wall includesopenings (pores or slits) with dimensions too large to restrictdiffusion of even the largest molecules. These openings serve asnon-diffusion-limiting exit ports within the capsule wall and protectinternal components of device 10. In this embodiment, nanopore membrane26 is seated next to or near an opening in the wall of a tubular inset,the outer diameter of which is similar to the inner diameter of thecapsule. The inset tube segment is inserted into the interior space ofthe capsule and the nanopore membrane array is positioned beneath theopenings. The inset tube is configured with o-rings, which serve to sealthe space between the capsule openings and the surface of the nanoporemembrane. As with the embodiment detailed above, nanopore membrane 26again serves as the only constraint to free passage into the reservoirfilled with the therapeutic agent from the environment outside of orexternal to the capsule.

As stated, drug delivery device 10 is designed to contain a volume oftherapeutic agent or other substance in reservoir 22. In an exemplaryembodiment, the capacity of the drug reservoir is approximately 500 μL.The therapeutic agent may be present in a variety of forms including,but not limited to, an aqueous solution, an aqueous crystallinesuspension (slurry), a micronized suspension. These solutions orsuspensions may be formed within reservoir 22 immediately beforeimplantation of device 10 by hydration of a dry form of the therapeuticagent, or by hydration after implantation of device 10 by inflow of abiological fluid from the medium surrounding the device.

Suitable therapeutic agents include biologically active macromoleculessuch as peptides, protein drugs, or a polynucleic acids. Suitablepeptides or protein biopharamceuticals include: hormones, hormoneagonists, hormone antagonists, growth factors such as CSF. EPO, andgrowth hormone, cytokines such as the interleukins, immune modulatorssuch as interferon gamma and interferon beta, anti-infectives such asinterferon alpha 2 b, anti-inflammatories, immunesuppressant/anti-rejection drugs, antibodies, anti-arthritic drugs, andanti-tumor agents. Suitable polynucleic acids include: DNA, RNA, plasmidmolecules, antisense DNA, and ribozymes. Small molecular weightmolecules are also compatible with the present invention. Suitable smallmolecular weight molecules include, but are not limited to, painmedications or anti-psychotic agents.

In an exemplary embodiment, drug formulations are filled into device 10as a liquid solution; however, if necessary more highly concentratedforms including a slurry consisting of an insoluble drug suspension, drypowders or centrifuged pellets may be used. For certain unstable drugs,such as certain proteins and peptides, the drug may be co-formulatedwith a bulking agent such as lactose and loaded into the reservoir ofthe device as a liquid. After loading the drug/lactose solution may befreeze-dried in situ, providing a dry storage form for the drug withinthe device reservoir. Excipients may be added to improve drug loadingsuch as polymers, ion-exchange beads, affinity matricis, cyclodextrins,surfactants and the like. Co-solvents such as ethanol and DMSO may alsobe employed for the purpose of loading additional drug.

Preferably, stabilizers co-formulated with therapeutic agent containedwithin reservoir comprise water miscible solvents, or polymers ofsufficiently high molecular weight and/or shape as to be retained withincapsule reservoir because they cannot pass out through nanopore membrane26. Suitable stabilizers include, but are not limited to carbohydrates,dextrans, polyvinyl pyrrolidone, gum arabic, polyethylene glycol,albumin, dendritic polymers, cross-linked polymer matrix, andsurfactants.

II. Nanopore Membrane

Prior art nanopore membrane designs (referred to as the “Berkeley”design) include arrays of C-shaped annular, or rectangular channels withprecise dimensions which in their smallest aspect, are in the 4-100 nmrange. C-shaped annular pores which are approximately 9 μm incircumference with 2 anchor regions result in a low porosity (0.26%) fora typical 25 nm pore membrane. The minimum reported pore size producedby the Berkeley design was 18 nm. In the original Berkeley design, thepores were produced in a 1.4 by 3.4 mm area (4.76 mm²) within a 6 by 8mm, rectangular-shaped solid silicon die region. U.S. Pat. Nos.5,651,900, 5,770,076 and 5,849,486 disclose microfabrication techniquescompatible with the present invention and are hereby incorporated byreference in their entirety.

One aspect of the present invention is based on the unexpected findingthat such membranes can be tailored to control the rate of release ofdrugs from an implantable device. By precisely fabricating pores withdimensions that are selected to range from a size similar to that of thedrug molecule to several times the size of the drug, (e.g.,approximating 1 to 5 times the Stoke's diameter of a drug molecule),nanopore membranes can be used to control diffusion at a molecular levelby providing release kinetics which approach zero-order. Zero-orderrefers to a rate of diffusion that does not change as the concentrationgradient decays. To achieve suitably sized pores, the nanopore membraneincludes a microfabricated array of parallel channels, wherein thesmallest aspect of the channels is selected to provide a constant rateof release of the drug or other substance.

The exemplary embodiment of the present invention was created using thebasic microfabrication approach pioneered at Berkeley. However, as shownin FIGS. 1 and 2, the exemplary embodiment includes a novelparallel-pore design. This pore design comprises a series of parallel,rectangular-shaped, 45 μm long channels that are separated lengthwise by10 μm anchor regions (FIG. 2). The width of these rectangular channelsis typically selected at fixed values in the 2 to 100 nm range. Using a2 pore spacing in the width direction, a high membrane porosity of 1% isachieved for the same 25 nm pore size. Pore sizes down to 4 nm have beenproduced by this approach. The die layout is also changed relative tothe Berkeley design. A similar large-die design is produced with 6 by 8mm overall dimensions, but the pore area is increased to 2 by 3.5 mm (7mm²). A smaller membrane die is also produced. In this case, themembrane area is 1 by 2 mm (2 mm²) within a 3 by 4 mm solid die area.

For the purposes of the present invention, microfabrication facilitiesavailable at The Case Western Reserve University (CWRU) in Cleveland,Ohio were used to produce several batches of the new nanopore membranedesign. This procedure uses a combination of photolithography anddeposition/selective removal of sacrificial layers. The pore size itselfis determined by the deposition and selective removal of a sacrificiallayer (silicon dioxide) sandwiched between two silicon structuralmembers as in FIG. 3. Using this approach, a selected array of parallelmembrane pores, with a specific pore size, density, and path lengthparameter, is fabricated as desired.

The method for fabricating nanometer-sized (nanometer wide and 45 umlong) channels consists of three basic steps: (1) surface micromachiningof the nanochannels in a thin film on the top of a silicon wafer bydepositing a nanometer-thick sacrificial oxide layer between two siliconstructures; (2) forming the membrane by etching away the bulk of thesilicon wafer underneath the thin-film structure, and (3) etching thesacrificial layer to form nanometer-sized pores in the membrane.Silicon, poly silicon, a combination of silicon materials, polymer, andco-polymer are all useful for fabricating the nanopore membrane usingthe techniques described.

The first major process step (see FIG. 1) involves photo orplasma-etching continuous channels in the silicon substrate over itsentire surface to define the overall pore shape 100. The etched channelsare typically 2 μm wide separated by 2 μm and the depth of thesechannels defines the membrane thickness of 5 μm. After forming thechannels, the next major step involves growing a sacrificial thermaloxide layer over the entire wafer surface including the surface area ofthe channels 101. The oxide layer thickness defines the pore size in thefinal membrane. Precise control is needed at this step to assure auniform, known pore size. Thermal oxidation gives pore size (width)control with sub-nanometer resolution across the entire 4-inch siliconwafer. Etching away the thermal oxide in selected areas of the channelforms anchor regions. The distance between these anchor regions definesthe pore height (45 μm).

A polysilicon structural layer is then deposited over the channels 102and planarized to allow access to the nanopore channels from the frontface of the array 103. Following a boron-doping step to create a KOHetch stop 105, a silicon nitride protective layer is deposited andwindows are opened in this layer on the backside of the wafer 104. Thebulk silicon is removed through these etch windows up to the etch stoplayer of the array using KOH 106. This bulk etch defines the membranearea and dye shape. The structure is then released by etching theprotective nitride layers and the sacrificial oxide layers in aconcentrated HF bath 107.

The major structural component of the nanopore membranes created by thisapproach is illustrated diagrammatically in FIG. 2. An array of parallelrectangular channels is formed 201 between alternating silicon 200 andpolysilicon 202 structural layers following the removal of thesacrificial oxide layer. Anchor regions 203 which lie between thechannel arrays provide mechanical strength.

Two batches of membranes were fabricated for the present invention.Table 1 provides a list of the number of membrane dies produced. In thistable, the batch number is listed along with the number of both largeand small dies produced with a given pore size. In all cases, the poredepth is 4.5 μm. Note that both large and small-sized dies were producedand that these membranes have a variety of pore sizes in the 2 to 100 nmrange.

TABLE 1 Nanopore Membrane Dies Batch Nominal Pore Number of Dies NumberSize, nm Small Large 2 4 121 53 1 7 391 263 1 13 340 268 1 20 92 69 1 27116 79 1 47 72 24 1 49 187 136

Other embodiments of the nanopore membrane of the present inventionincorporate the following characteristics, which presumably improve orenhance the durability of the membrane: (i) a multi-directional porepattern rather than orienting all the pores parallel in only onedirection which tends to increase stress; (ii) reduced membrane area;(iii) incorporation of membrane material that is stronger than siliconsuch as silicon carbide; and (iv) the additional of a hexagonalhoneycomb support structure to the surface of the membrane thatincreases its effective thickness without changing the pore lengthmembrane thickness.

III. Osmotic Engine

Release of a therapeutic agent from the drug delivery device of thepresent invention is based on a constrained diffusion mechanism. Thatis, the rate of diffusion of the therapeutic agent through a staticaqueous phase filling the channels of the nanopore membrane isrestricted by the width of the channels. It is possible to facilitate or“boost” diffusion through such channels by providing a net movement orflux of fluid in an outward direction relative to the drug reservoir,through the nanopore channels. Such a fluid flux may be provided byequipping the device with an osmotic engine. Therefore, in an alternateembodiment of the present invention the diffusional release mechanism ofdrug delivery device 10 is facilitated by incorporation of an osmoticengine into the device described above.

With reference to FIG. 3C, this osmotic engine includes (i) asemi-permeable membrane 30 incorporated into a portion of the wall ofthe capsule, wherein the membrane is permeable to the passage of fluidpresent in the environment external to capsule 12 and substantiallyimpermeable to the passage of the therapeutic agent in reservoir 22;(ii) an osmotically active agent admixed with the therapeutic agent,wherein the osmotically active agent is of sufficient molecular weightas to be restricted from passing through both the semi-permeablemembrane and the nanopore membrane; and (iii) a net flux of waterentering the reservoir 22 from the external environment throughsemi-permeable membrane 30 and exiting through the nanopore membrane.

The osmotic engine consists of two basic elements. The first is asemi-permeable membrane introduced into a portion of the containmentwall of the capsule. For example, one (or both) of the solid end capsdepicted in FIG. 1 may be substituted with a plug made of asemi-permeable material such as cellulose acetate (see FIG. 9). Such amembrane permits entry of fluid from the medium or environmentsurrounding device 10, but prevents the release of solutes containedwithin the reservoir 22. Listings of suitable semi-permeable membranematerials and permeability-modifying additives are found in U.S. Pat.Nos. 4,077,407 and 4,874,388, and are available from Eastman ChemicalCompany (Kingsport, Tenn.). Precast cellulose acetate and othercellulose ester membrane sheet stock are available from Polymer ExtrudedProducts (Newark N.J.) and LAM Technologies (Berlin, Conn.).

The second component of the osmotic pump is an osmotically active agent,or osmagent, admixed with the therapeutic agent. The osmagent is ofsufficient molecular weight as to be restricted from passage throughboth the semi-permeable membrane and the nanopore membrane and thus issubstantially retained within reservoir 22. The osmagent preferred forbiologically labile protein therapeutics comprises a water-miscible,liquid polymer in which a dry, micronized form of the drug is suspended.Thus, the osmagent may be linear or branched water miscible polymers ofMW 5000 to several million Daltons, including but not limited to,dextrans, maltodextrins, albumin, starches, polysucrose, Ficoll,dendritic polymers, hyaluronic acid, polyvinyl alcohol, polyethyleneoxide-polypropylene oxide copolymers, polyethylene glycol, andpolyvinylpyrrolidone. The osmagent may also be fluid polymers up to50:50 mixture by weight including but not limited to: Pluracol V-10, amember of the UCON lubricant series, a member of the Pluronic surfactantseries, or a member of the Tetronic surfactant series.

The present invention is intended to release encapsulated drug atphysiologic temperatures and over time frames ranging from weeks tomonths. Some clinically important biologically active proteintherapeutics may be unstable in aqueous solution under such conditions.While the therapeutic agent may be formulated as an aqueous orcrystalline suspension which may or may not include stabilizers such asthose described above, a preferred osmotic engine design forbiologically labile protein therapeutics is a dry, micronized form ofthe drug suspended in the anhydrous, water-miscible, liquid polymerosmagent. A drug used with this embodiment may be micronized by anynumber of methods known to the art such as spray drying, freeze dryingand/or milling. The micronized drug and osmagent are mixed to create asuspension that is filled into the device reservoir. The osmagent may bemixed with a small fraction of water (for example, up to 50:50 byweight) if desired, for example, to change viscosity and/or to prime thesystem for diffusion-based operation.

The osmagent creates a colloid osmotic pressure within reservoir 22,attracting water from the external medium. Capsule 12 is not expandableand thus the volume of the reservoir is held constant. Water enteringthe device through semi-permeable membrane 30 under the influence of theosmagent is thus forced to exit through the nanopore channels. The rateof flux of water is controllable by (i) the features of thesemi-permeable membrane including the material used, additives, inherentwater permeability, surface area, and thickness, and (ii) the nature ofthe osmotically active polymer including the molecular weight,concentration, and osmotic pressure. Methods to calculate permeationrates through semi-permeable membranes under selected conditions aredisclosed in U.S. Pat. No. 4,077,407 which is hereby incorporated byreference in its entirety.

Fluid flux provided by such an osmotic engine is used to boost efflux ofdrug molecules from the drug delivery device when necessary to achieve adesired release profile (i.e., zero-order) and/or desired durations ofrelease. The fluid flux is also intended to maintain the patency of thenanopore fluid channels by inhibiting fouling or clogging of thenanopore channel walls and end-openings. The fluid flux also provides ameans to slowly dissolve therapeutic agents which have been loaded intothe device reservoir in an insoluble, partially soluble or dry form.

In a preferred configuration, following implantation, external fluidenters the device through semi-permeable membrane 30, under theinfluence of the water-attracting properties of the polymer osmagent,and exits through the nanopore membrane while the volume of thereservoir remains constant. Entry of water into device 10 slowlysolubilizes the exposed surfaces of the micronized drug particles, andsmall amounts of drug become entrained as a solute in the fluid flow.The amount of water passing through the device at any given time islimited and thus the protein enters solution gradually over time. Sincethe dry particles dissolve slowly, the stability of the micronizedprotein drug is preserved. Release of drug molecules from the device isbased on diffusion through the fluid phase and is controlled by waterflux through the reservoir, the rate of solvation of the drug into thewater phase, and by selection of the proper nanopore membrane dimensions(as detailed below). Combined, these features provide a high degree ofcontrol over release kinetics.

The osmotic influx of fluid provided by the osmotic engine feature ofthe present invention also allows for in vitro or in situ activation ofthe device. For example, during manufacturing, drug delivery device 10may loaded with a dry form of a protein drug-osmagent admixture topreserve biological activity of the protein drug during distribution andstorage. At the time of use, the device may be placed in a sterileaqueous solution, such as PBS or D5W, for a period of time sufficientfor entry of a portion of the external aqueous medium into the devicereservoir through the semi-permeable membrane. The fluid entering thereservoir dissolves or partially dissolves the protein drug and osmagentand prime the nanopore channels with fluid, thus providing the aqueouspathways required for the constrained diffusion mechanism to operate.The device would then be rinsed with sterile PBS or D5W and implanted.Similarly, a device loaded with a partially dissolved or dry therapeuticagent-osmagent admixture may be implanted into a body compartment suchas under the skin. The external fluid medium would enter the devicereservoir through the semi-permeable membrane, hydrating thedrug-osmagent admixture and priming the nanopore channels. A slightdelay in drug efflux is expected as priming takes place.

In summary, the combination of the physical constraints imposed on thediffusion of drug molecules provided by the nanopore membrane geometry,plus the osmotically driven flux of fluid through the device, provides anovel means of tailoring the release kinetics of selected drugs toachieve optimal therapeutic outcomes.

IV. Examples

Controlling the rate of drug diffusion through nanopore membrane 26 isan important aspect of the present invention. Previous researchinvolving nanopore membranes has not recognized the relationship betweennanopore membrane geometry and the phenomenon of constrained moleculardiffusion. Furthermore, previous research has not suggested the benefitsof combining constrained diffusion with an osmotic engine as detailedabove. Thus, diffusion studies have been conducted for the presentinvention to evaluate the applicability of nanopore membranes toconstrained molecular diffusion for a small molecule (glucose), a largemolecule (albumin), and lysozyme. These studies, detailed in theexamples below, provide a demonstration of the ability of nanoporemembranes to constrain rates of diffusion over a broad range ofmolecular weights.

Example 1 Glucose Diffusion

Nanopore membranes with channel widths ranging from of 7 nm, to 27 nmwere microfabricated from silicon materials according the methodsdescribed above. These membranes were placed either in an implant devicefitted with nanopore membranes as described above, or placed in a COSARTdiffusion cell. These diffusion cells consist of two reservoirs with thenanopore membrane sandwiched between. The membrane is sealed by the useof appropriately sized o-rings. The lower reservoir is equipped withinlet and outlet ports through which buffer (such as phosphate bufferedsaline, PBS) is slowly pumped (using either an HPLC pump or syringepump). The lower flow cell also contains a stir bar for mixing. Asolution of glucose is added to the upper reservoir and fractions of thefluid exiting the lower reservoir collected using a fraction collectoror by physically moving the outlet tubing from one empty test tube toanother at preset time intervals. The glucose concentration in thefractions is measured using a either a standard glucose monitor orstandardized glucose-oxidase assay.

The cumulative diffusion of glucose through nanopore membranes of fourdifferent channel widths, 7 nm, 13 nm, 20 nm, and 27 nm is plotted inFIG. 4. The shapes of the curves for the 13 nm, 20 nm, and 27 nmmembranes indicates that the cumulative movement (flux) of glucosethrough the membrane is non-linear, and as a consequence the rate slowsrelative to the initial rate as the concentration gradient across themembranes decays. The rate of glucose diffusion through the 7 nmmembrane, however, is linear or zero-order. This pattern is unexpectedsince the diffusion gradient is decaying with time and thus the drivingforce for diffusion would be expected to decay as well.

The glucose diffusion rate (expressed as mg/day) through nanoporemembranes of 7 nm, 13 nm, and 49 nm widths as a function of initialglucose concentration is shown in FIG. 5. For the 13 nm and 49 nmmembranes, the rates of glucose flux increases with increasingconcentration. In the case of the 7 nm membrane, however, a flux rate of1.2 mg/day is maintained across a wide range of initial glucoseconcentrations (i.e., 165-660 mg/mL). These data indicate that glucosediffusion through nanopore membranes with widths of about 7 nm and belowis basically zero-order and is not influenced by the magnitude of theconcentration gradient across the membrane.

Example 2 Albumin Diffusion

For these experiments, Radiolabeled (¹²⁵I) bovine serum albumin (BSA)was obtained from ICN Biochemicals; Unlabeled BSA was obtained fromSigma Chemical Corp., and microfabricated silicon nanoporous membraneswere fabricated as described above.

Implants fitted with nanopore membranes were assembled and loaded with¹²⁵I BSA. Implants were sterilized prior to loading by eitherautoclaving or chemically sterilizing by soaking in a standardphosphate-buffered saline solution (PBS) containing 0.2% (w/v) sodiumazide. The implant formulation was generated by diluting the appropriateamount of ¹²⁵I-labeled BSA (determined from the specific activityreported by the manufacturer for each lot obtained) with a 25 mg per mLstock solution of unlabeled BSA. The solution was filter sterilized. Thefinal formulation solution contained BSA at a concentration of 5 mg permL and a ¹²⁵I-labeled BSA concentration of 60-65 μCi per 300 μL, thecurrent approximate loading capacity of each implant.

After aseptically loading the implants with the BSA formulation, eachimplant was washed by dipping into 25 mL PBS and transferred to a fresh25 mL PBS, both contained in sterile 50 mL conical tubes. Any residualair bubbles, trapped in the grate opening above the membrane area, weredislodged by moving a stream of liquid containing PBS over the gratingwith a 1 mL syringe containing a 27 gauge luer-lock needle. A 100 μLaliquot was removed once a day for 3 to 6 days or until the expectedconstant rate of diffusion was achieved. The aliquot was transferred toa 5 mL polypropylene serum tube and counted on a Packard Gamma Counter.

For each treatment group, three, ten to twelve-week old female SpragueDawley rats were anesthetized with isoflurane. After shaving andsterilizing the implant site, a 1.5 cm incision was made parallel to and1 cm from the backbone through the skin layer midway on the animals'flank. A pocket was made for the implant under the skin by opening andclosing a pair of hemostats to the front and back of the incision. Theincision “pocket” was made in this way to ensure that the ends of theimplant did migrate to a position over the incision, limiting thepossibility for the ends of the implant to push out through theincision. The implant was rinsed in 2 changes of PBS; the grate of theimplant was wetted with 10 uL of PBS. A small amount (100 μL) of PBS wasintroduced into the pockets for wetting purposes and the implants wereinserted into the incision pocket and centered under the incision site.The incision was sutured and the animal was allowed to recover.

A small plastic ring was used to orient the implant membrane downwardand in the proper direction relative to the animals' flesh and skin.Release in the proper direction insured that the implant diffusion wasnot impeded in any way by physical blockage.

For pharmacokinetic analysis, blood sample were obtained. Animals wereanesthetized with Isoflurane, the tails warmed for 2 minutes, and 0.2 mlsamples of blood were collected from the lateral tail veins. The bloodwas allowed to clot for 45 minutes, and then spun at 10,000 G for 1minute, and then a 50 μL, aliquot removed for gamma counting on aPackard Gamma Counter.

The results of these in vivo experiments, shown in FIG. 7, represent thetotal calculated amount of BSA measured in the vascular compartment (inμg per mL) of each of the three treatment groups (n=3): subcutaneousinjection, or NanoGATE devices designed to deliver 7 μg/day or 15μg/day. In the case of the implant groups, a two-phase pattern is seenfor BSA clearance from the blood stream. The first phase, represented bydata collected between days 2 through 10, shows a rapidly decreasingslope that closely resembles the slope obtained with a 125 [μgsubcutaneous (SubQ) injection. This rapid initial clearance of BSA fromthe bloodstream is very likely due to the extravasation of BSA fromcapillary beds into the interstitial fluid compartment and eventuallythe lymphatic system. Equilibration of exogenously administered BSAbetween the blood and interstitial compartments is a well knownphenomenon.

The second phase of BSA serum levels shown in FIG. 7 is represented bydata collected between days 10 through 50. During this phase, BSA isslowly cleared from the animals' blood. These data suggest thatphysiological equilibrium or steady-state is achieved suggesting thatthe rate of BSA release from the NanoGATE implant is balanced withclearance from the blood stream. This pattern stands in stark contrastto the rapid clearance of the onetime, 125 μg bolus dose of BSAadministered via subcutaneous injection.

Example 3 Lysozyme Diffusion

Lysozyme, an enzyme with a molecular weight intermediate between that ofglucose and albumin (MW 12,000 Daltons) was used to measure diffusionkinetics through nanopore membranes both in vitro and in vivo.Radiolabeled (¹²⁵-I) lysozyme was obtained from ICN Biochemicals.Unlabeled lysozyme was obtained from Sigma Chemical Corp. Radiolabeledlysozyme was admixed with unlabeled lysozyme and loaded into implants asdescribed above for albumin. In vitro release rates for three sizemembranes are plotted in FIG. 8. Lysozyme release rates are zero-orderfor membranes pore sizes of 13 nm and 49 nm. Release of lysozyme throughthe 7 nm membrane implant was virtually undetectable over the time frametested (5 days). Additional in vitro release testing revealed that, likeglucose, at a pore size of 13 nm or less, flux of lysozyme through thenanopore membrane is independent of initial concentration (data notshown).

Blood levels of lysozyme in groups of three rats given either a singlesubcutaneous bolus of 80 μg or after implantation of a nanopore implantdevice designed to release lysozyme at a rate of 17 μg/day are shown inFIG. 9. Lysozyme given subcutaneously is cleared rapidly from the bloodstream of the animals with an apparent half-life of less than severaldays. In contrast, blood levels in animals in the implant group weremaintained for a period of several months.

V. Method of Use

The present invention is useful for delivering a therapeutic agent tothe body in a controlled manner, according to the following method.First therapeutic agent is loaded into the device, and then the deviceis surgically implanted into the patient. As detailed above, an aqueoussolution or suspension of the therapeutic agent may be formed inreservoir 22 immediately before implantation by means of hydrating a dryform of the therapeutic agent. An aqueous solution or suspension of thetherapeutic agent may also be formed in the reservoir after implantationby introducing a biological fluid derived from the medium surroundingthe device into the device.

While the drug delivery device 10 is in use, the rate of release of thetherapeutic agent is substantially zero-order and is typicallyindependent of the concentration of the therapeutic agent on theinterior of the capsule. The rate of release of the therapeutic agent issubstantially constant for a period of several weeks to several monthsfollowing surgical implantation of the device.

While the above description contains many specificities, these shouldnot be construed as limitations on the scope of the invention, butrather as exemplification of preferred embodiments. Numerous othervariations of the present invention are possible, and it is not intendedherein to mention all of the possible equivalent forms or ramificationsof this invention.

What is claimed:
 1. A method for delivering a therapeutic agent,comprising: providing a device comprising a fluid impermeable walldefining a reservoir that contains a therapeutic agent with moleculardimensions; said capsule further comprising an exit port incommunication with the reservoir and a nanopore membrane incommunication with the exit port, wherein said nanopore membranecomprises an array of annular pores, each pore in the array having adiameter of between about 2-100 nm and dimensioned to be between about1-5 times the molecular dimensions of the therapeutic agent, and whereinsaid therapeutic agent is released by diffusion across said array ofannular pores at a rate that approaches zero-order for a period of aboutseveral weeks to about six months, said device being indicated forimplantation.
 2. The method of claim 1, wherein said device is indicatedfor subcutaneous implantation.
 3. The method of claim 1, wherein saidarray of annular pores is an array of parallel annular pores.
 4. Themethod of claim 1, wherein the therapeutic agent is dispersed withinsaid reservoir as an aqueous solution or aqueous suspension.
 5. Themethod of claim 1, wherein the therapeutic agent is in dry form.
 6. Themethod of claim 1, wherein the therapeutic agent is a peptide.
 7. Themethod of claim 1, wherein the therapeutic agent is a protein.
 8. Themethod of claim 1, wherein the protein is growth hormone.
 9. The methodof claim 1, wherein the therapeutic agent is an anti-psychotic agent oran agent for pain.
 10. The method of claim 1, wherein the device iscomprised of titanium.
 11. The method of claim 1, wherein the capsule iscylindrical.